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pressure drops between arterialand venularchannels [26]. Theircombination of dened inlet and outlet branches as provided bythe microuidic device with the formation of a genuine capillarynetwork within the device enabled perfused vascular networks invitro, but in its current form is not useful for the assembly of 3Dtissue or implantation into living organisms. With respect to invivo application of articial tissue, it has a been observed in earlystudies by Tremblayet al.[27] and has also been acknowledged in

recent studies performed, for example, by Nunes et al. [28] thattissue structures containing fragments of capillaries connectedmuch faster to the native vascular system of a host upon implant-ation than tissue without such vascular prestructures. Yet, suchconstructs, on the other hand, do not provide continuous tubularvessel systemsin vitroand, therefore, they cannot be used for thesupply of surrounding cells as needed for 3D TE.

Thus, the essential requirements for fabrication of perfusablevascularized tissue seem to be both, to have dened inlet andoutlet branches and to provide a capillary-like network with bio-mimetic vessel geometry for undisturbeduid dynamics [29].

Therefore, additional approaches provide ready-made tubes ortubular networks as a 3D matrix for supplying cells in 3D culture.These may be either based on the reuse of biologically derived,

decellularized vessel systems [30] or synthetically manufacturedtubular scaffolds [31]. To develop articial structures that performas well as natural ones, we need fabrication processes that do notset any limits to the generation of structures and shapes, andmaterials that allow for tailoring of their physical, chemical andbiological properties. Bioprinting technology, which applies free-form fabrication methods to deposit scaffold materials and cells toform digitally dened 3D structures [32], constitutes a basicallynovel approach to approach this so far unsolved issue: There arevarious types of 3D manufacturing techniques addressing a broadrange of structure sizes and several have already been used forbiofabrication purposes as well.

THE POTENTIAL OF BIOPRINTING TECHNIQUES

Bioprinting means the generation of bioengineered structures byadditive manufacturing of biological and biologically relevantmaterials with the use of computer-aided transfer and build-upprocesses [3335]. A variety of solid freeform fabrication techni-ques are available. They can be divided into laser-based methods,printer-based methods and nozzle-based methods.

Nozzle-based systems use pressure-assisted syringes to depositcontinuous strands of materials. They typically address structuresizes in the centimetre range with resolutions of several hundredmicrons [36]. Printer-based systems include thermal and piezo-electric inkjet printing. These drop-on-demand systems generatesmall droplets of low viscosity bioink [37]. Typical resolutions are

in the range of 85

300 m and can be used for generation ofmillimetre to centimetre constructs [35]. Laser-based systems in-cluding stereolithography and multiphoton polymerization as wellas digital light processing (DLP), a method based on digital micro-mirror devices [38], use light for site-selective curing of photo-sensitive prepolymers in a bath with resolutions down to thesub-micron range [39].

Three parallel bioprinting-based approaches are followedtowards the aim of the generation of articial blood vessels: (i)generation of bulk matrices with integrated channels as perfusablematrices, (ii) cell patterning into line structures for self-assembly ofinterconnected vessel systems and (iii) generation of free-standing

tubular structures, with and without encapsulated cells, whichmay serve as articial blood vessels (Table1).

PERFUSABLE MATRICES

The most straightforward approach to perfusable tissue mightbe the generation of a network of interconnected channels within

the tissue matrix. Such channels may be used as supply system forcells within the surrounding matrix and may additionally beseeded with ECs. Early works used moulds for preparing sacricialstructures in order to fabricate microuidic networks, which thenallowed the transport of macromolecules into surrounding hydro-gels under low driving pressure differences. Such studies resultedin zones of high cell viability up to 200 m away from the perfusedchannels and thus demonstrate the effectiveness of the strategy.Yet, the soft lithographic techniques that were applied for channelgeneration require multiple processing steps and are limited toplanar networks and stacks of planar networks [6567]. Anotherapproach presented by Bellanet al. used melt-spun bres as a sac-ricial mould and thus achieved a 3D channel network, yet inrandom spatial orientation [40].

For rapid casting of a de

ned pattern of vascular channels,Miller et al. developed a sacricial material based on carbohy-drates, which they printed in laments using a syringe and aRepRap Mendel 3D printer [41]. They used mixtures out ofglucose, sucrose and dextran (86 kDa) as liquid glass and pro-duced self-supporting perpendicular lattices as well as curvedla-ments and Y-junctions with bre diameters in the range of 150750 m (Fig.1). Such glass lattices were coated with thin layers ofpoly(lactid-co-glycolid) (PLGA) and subsequently encapsulated inthe presence of living cells using a wide range of natural and syn-thetic matrix materials such as brin, agarose and Matrigel, andpoly(ethylene glycol) (PEG)-based materials. After cross-linking ofthe matrix, the carbohydrate support was dissolved in aqueousmedia. Thereby, the PLGA layer prevented carbohydrate solution

from owing through the cell-laden hydrogel. The resulting chan-nels allowed non-leaking perfusion of human blood. ECs wereseeded into the channels and lined the walls of the completenetwork. The authors observed sprouting of new capillaries fromthe channel walls into the surrounding brin gels which had beenladen with 10T1/2 cells. In agarose gels and even in gels madefrom synthetic PEG hydrogels, which were modied with the RGDcell recognition peptide, various cell types were live and active atthe gel perimeter and approximately 1 mm around the perfusionchannels. For comparison, within gels without channels, cells wereonly active at the gel slab perimeter but not at the gel core.

Leeet al. dispensed heated gelatin solution for the generationof sacricial elements and ice-cooled collagen hydrogel precursorat a pH of 4.5 to produce collagen layers, which were then cross-

linked by spraying NaHCO3 solution on top [43]. Dermal bro-blasts grown in such collagen scaffolds containing uidic channelsalso showed signicantly elevated cell viability compared with theones without any channels, proving again the effectiveness of in-tegrating vascular structures for 3D tissue reconstruction.

In their approach to fabricate a free-form 3D channel networkwithin a hydrogel, Wu et al. used a reservoir of photo-cross-linkable, acrylated Pluronic F127 (PF127, 2025% w/w) anddeposited sacricial lament networks from nonacrylated PF127directly into such reservoirs [42]. PL127 is an interesting materialfor manufacturing purposes because it shows unusual properties,i.e., it is liquid at 4C and solidies when warmed. It is one of the

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Table 1: Overview over biofabrication methods and materials used in present approaches for the generation of vascular structures

Bioprinting technique Building material/bulk hydrogel Sacrificial material/supportivematerial

Cell type

Perfusable matricesNozzle-based:

Modified cotton candy machineEnzymatically cross-linked gelatin Shellac

Nozzle-based:RepRap Mendel 3D printer

PEG diacrylate + acrylate-PEG-RGDS; Matrigel;agarose; alginate; fibrin

Carbohydrate glass(glucose + sucrose + dextran)

HUVEC, 10T1/2

Nozzle-based Acrylated Pluronic F127 Pluronic F127 Nozzle-based Collagen Gelatin Primary HDFs Nozzle-based Mixtures of gelatin, a lginate , chitosan, f ibrinogen, HA Rat primary hep

ADSCsCell patterning for autonomous vessel formation

Laser-based:BioLP

HUVEC, HUVSM

Laser-based:BioLP

Alginate, Matrigel, fibrin Rabbit carcinomB16, HUVEC cEathy926

Drop-on-demand:HP Deskjet 500

Fibrin HMVEC

Nozzle-based:BioAssembly Tool

Collagen RFMF

Free-standing tubular structuresLaser-based:

Two-photon polymerization

,-Polytetrahydrofuranether-diacrylate

Laser-based:Digital light processing

Formulations based on urethan-diarylate, hydroxylethylacrylate as reactive diluent, ethylene glycolbisthioglycolate as the chain transfer agent

Drop-on-demand:Thermal inkjet, SEAjet

Alginate Ca solution + HA/PVA

Drop-on-demand:Thermal inkjet, SEAjet

Alginate Ca solution + PVA HeLa

Drop-on-demand:Thermal inkjet, HP697c

Alginate Alginate solution Rat SMC

Drop-on-demand:Piezoelectric MicroFab

MJ-ABL-01-120-6MX dispensehead

Alginate Ca solution NIH 3T3 fibrobl

Drop-on-demand:Custom-made, syringe + switchable

valve applied droplet volumeof 116 nl

Agarose Hydrophobic high-densityperfluorotributylamine C12F27N

Human MG-63 human MSC

Nozzle-based:Novogen MMX Bioprinter,

Organovo

NovoGel Mouse embryofibroblasts

Nozzle-based:Fab@Home

Thiol-modified HA + thiol-modified gelatin with goldnanoparticles or tetra-acrylated PEG as cross-linker

HA NIH 3T3 fibrobl

Nozzle-based:Fab@Home

Methacrylated HA + methacrylated gelatin HA Human hepatomHepG2 C3A


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few synthetic polymer materials that are approved for clinicalapplications by the American Food and Drug Association FDA[68]. However, the presented study does not discuss the integra-tion of cells into the constructs, and other studies revealed non-cytotoxicity of PF127 gels only up to 20%. [69,70] Therefore, weconsider the use of acrylated PF 127 as a cross-linkable matrix ma-terial to provide a proof of concept of omnidirectional lamentextrusion within a hydrogel precursor matrix rather than a real

matrix for 3D TE.Liet al. introduced bio-based material systems that were suitedfor vertical assembly of hollow channels without the use of a sup-portive sacricial material (Fig.2). The authors used different com-binations of mixed gelatin/alginate/chitosan/brinogen hydrogelsas building materials. Their printed structures were rst stabilizedby the solgel transition of the gelatin as the mixtures (37C) wereextruded from the syringe needle into a low-temperature fabrica-tion chamber (68C). The alginate, chitosan and brinogen werethen subsequently cross-linked using thrombin, CaCl2, Na5P3O10and glutardialdehyde [44].

CELL PATTERNING FOR SELF-ASSEMBLY OFINTERCONNECTED VESSEL SYSTEMS

The second approach to generate blood vessels in articial tissuedeals with the ability of ECs to organize into blood vessels autono-mously (angiogenesis). In vivo angiogenesis occurs as a responseto hypoxia, for example, in growing tissue when ECs sprout fromexisting blood vessels to form new capillaries by the curling of asingle or few cells [71]. While for in vitro TE, an interconnectedsupply system has to be instantaneously available, genuine vascu-lar structures may mature in parallel for the fast integration of theengineered tissue to the host tissue upon implantation. It hasbeen observed that cultivation of ECs that were randomly distribu-ted within 3D, growth factor-rich hydrogels, e.g. Matrigel, led to

formation of an undirected, polygonal cellular network [72]. Thus,it is expected that prestructuring of vascular cells may guide thedirection of lumen formation and growth into an interconnectedcapillary system. By the application of digital patterning techni-ques, directed cell deposition can be achieved and branchingstructures with biomimetic branching geometries may be designed.

Towards this aim, biological laser printing (BioLP), also referredto as laser-assisted bioprinting (LAB), allows for deposition of indi-vidual cells. This method is based on laser-induced forward trans-fer of biological materials and cells. Wu and Ringeisen [45] andGuillotinet al.[46] used LAB for positioning of individual humanumbilical vein endothelial cells (HUVECs). Wu and Ringeisendesigned dened branch and stem structures and observed thatdeposited HUVECs connected with each other within 1 day (Fig.3).

However, the connected branches built from HUVECs alone didnot last more than a few days. Stabilization of the cellular networkstructure for at least 9 days occurred only with the deposition of anadditional cell layer of human umbilical vein smooth muscle cells(HUVSMCs) on top of HUVECs. This conrms similar results fromother labs and indicates that assembly of ECs with additional vascu-lar cells, such as pericytes and smooth muscle cells, is crucial for thegeneration of stable, functional articial blood vessels and function-al tissue substitutes [73, 74].

While laser-based technologies allow for precise positioning ofindividual cells and thereby the generation of structures with reso-lution at the cellular level, technologies such as inkjet printing or

Table

1:

Continued

Bioprintingtechnique

Buildingmaterial/bulkhydrogel

Sacrificialmaterial/supportiv

e

material

Celltype

Channel

diameter

Ref.

Nozzle-based:

Fab@Home

Thiol-modifiedHA+TetraPAc13(cross-linker)

HA

NIH3T3fibroblasts

Diameter:500m

Wall:5

00m

[59]

Nozzle-based:

BioScaffolder3DFsystem,

SYS+ENG,Germany

Polycaprolactone

Diameter:24mm

Wall:1

mm

Height:20mm

Totals

ize:67428

mm

[60]

Nozzle-based:

BioScaffolder3DFsystemSYS+ENG,

Germany

Methacrylatedgelatin/gellangumhydrogel

Diameter:24mm

Wall:1

mm

Height:20mm

Totals

ize:67428

mm

[60]

Nozzle-based:

Multicellularspheroids

HUVSMCs,HDFs

Diameter:900m

Wall:3

00m

[61]

Nozzle-based:

Co-axialnozzleprinting

Alginate

Bovinecartilage

progenitorcells

Diameter:100600m

[6264]

ADSCs:adipose-derivedstromalcells;HDFs:humandermalfibroblasts;HUVECs:huma

numbilicalveinendothelialcells;10T1/2;HUVEC

s:humanumbilicalveinendothelialcells;HUVSMCs:humanumbilicalvein

smoothmusclecells;HMVECs:huma

nmicrovascularendothelialcells;RFMFs:ratfatm

icrovesselfragments;HUVSMCs:humanumbilicalveinsmoothmusclecells;MSCs:mesenchymalst

emcells;SMCs:smooth

musclecells;HA:hyaluronan;PEG:poly(ethyleneglycol);RGDS:arginineglycineasparticacidserine;BioLP:biologicallaserprinting;DO

D:drop-on-demand.



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dispensing enable the generation of structures in the micrometreand millimetre range. Drop-on-demand inkjet printing generatessingle droplets and can be used to process low viscosity hydrogelprecursors (10 mPa s and below) and cells. Such so-calledbioinks are cross-linked after the jetting to form a (bio-)articial

extracellular matrix (ECM) around the printed cells. Cui andBoland, for example from the Boland lab in Texas, which was oneof the cradles of bioprinting, built brin microstructures ladenwith human microvascular endothelial cells (HMVECs) using amodied thermal ofce inkjet printer [47]. They printed dropletsof thrombin solution and cells into a bath of brinogen on thesubstrate. The printed HMVECs were found to proliferate within 7days and align along the printed brin gel structure. With their ex-periment, Cui and Boland proved that thermal inkjet printing canbe used for cell and matrix patterning.

Besides isolated ECs, microvessel fragments can also be appliedas primary building blocks for autonomous vessel formation via

angiogenesis. Chang et al. prepared collagen structures in themillimetre range with encapsulated rat fat microvessel fragments(RFMFs) using a nozzle-based 3D bioprinting tool [48]. The nar-rowed extrusion of the relatively rigid RFMF within a liquid sus-pension (collagen solution) uniaxially aligned the RFMFs. After 7

days of culture in vitro and after subcutaneous implantation intoimmunocompromised mice, they found that numerous neo-vesselsspread from the parent RFMFs and were perfused. However, in themature network, the initial uniaxial alignment was lost. Conversely,when a uniaxial steel mesh frame was used as patterning cueinstead, which maintained tension across the 3D construct alsoduring implantation, the microvessel segments in the maturenetwork were oriented as intended. This experiment indicatesthat the angioarchitecture is profoundly inuenced by mechanicalforces that are derived from the surrounding tissue. Such experi-ence suggests that the initial patterning of cells or vascular frag-ments may not be sufcient when aiming at the generation of

Figure 1:Generation of perfusable channels represents one successful approach to improve cell supply within 3D hydrogel matrices. (A) Architectural design of a mul-tiscale lattice (green), (B) top view of the design shown in Aprinted in carbohydrate glass, (C) multilayered lattices were fabricated in minutes with precise lateral andaxial positioning, (D) endothelial cells formed single and multicellular sprouts (arrowheads) from patterned vasculature, as seen in az-stack (optical thickness = 200 m)from within the gel (z-position = 300 m), (E) serialy-junctions and curved laments were also fabricated, (Fand G) primary rat hepatocytes and stabilizing stromalbroblasts in agarose gels (F) slab gel, (G) channelled gel after 8 days of culture were stained with a uorescent live/dead assay (green: calcein AM; red: ethidiumhomodimer): Cells survived at the gel perimeter and near perfused channels, and survival decayed deeper in the gels. Scale bars are 1 mm. (Reproduced and slightlymodied from Milleret al.[41], with the permission of Nature Publishing Group.)

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vascular networks with tailored architectures such as denedinlet and outlet branches. A tailored environment providing ad-equate stimuli such as perfusion or appropriate stiffness might benecessary.

FREE-STANDING TUBULAR STRUCTURES

The technically most demanding approach to articial vascularsystems deals with fabrication of freestanding tubular structures.Such structures can then serve as articial blood vessels in 3Dtissue models and may also reveal some potential for small-diameter blood vessel medical grafts in the future. Recent studiestackle both paths, the preparation of tubular scaffolds withoutcells and the preparation of tubular structures, which are com-posed out of cellmatrix compounds or cell aggregates.

To enable the manufacture of tubular structures, i.e. structuresof very high aspect ratio, classically, photo-cross-linking-basedadditive techniques circumvent the problem of structure supportby selectively cross-linking the precursor polymer layer by layerwithin a reservoir of the polymer resin.

As part of the Fraunhofer BioRap platform, which is dedicated

to the development of additive assembly technologies forbiological applications, Engelhardt et al. and Meyer et al. usedstereolithography and two-photon polymerization (TPP) to fabri-cate tubular structures and bifurcated tubes out of photo-cross-linkable synthetic polymers and biopolymers. Meyer et al. pre-sented homologous series of photo-polymerizable ,-polytetrahydrofuranether (PTHF)-diacrylate resins, which afterphoto-curing cover Youngs Modulus in the range from 8 to 28MPa [49]. After appropriate rinsing, extracts of the cross-linkedPTHF polymers showed non-cytotoxic characteristics. In anotherwork, the authors introduced resins, which were also tailored forinkjet-printing applications and could be photo-cross-linked into

Figure 2:The layer-by-layer assembly of hydrogel mixtures combining the gelling behaviour of gelatin and cross-linking of additional biopolymers such as alginate orbrinogen enabled generation of vertical channels without the use of a support material. Dispensing of hydrogels with two different types of cells was demonstrated.Red, tube-like part: Adipose-derived stem cells from rat within gelatin/alginate/brinogen in DMEM/F12 cell culture media. White, meshwork part: Rat hepatocyteswithin gelatin/alginate/chitosan in PB; (Aand B) the control of salivation by improved processing intervals; (CE) the layer-by-layer fabrication process. (Reproducedfrom Liet al. [44], with the permission of Sage Publications.)

Figure 3:Positioning of single cells was demonstrated by application of laser-assisted bioprinting (LAB): Printed HUVEC double branch structure (parallel), 1day post-printing. The HUVECs from the two different branches reached acrossand networked. (Reproduced from Wu et al. [45], with the permission of IOPPublishing.)



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carboxy-terminated polyacrylate networks [31]. They observed thatcell adhesion and proliferation at the carboxy-functional surfaceof such synthetic materials were improved when the surfaces weremore hydrophilic.

Applying TPP, Engelhardtet al. generated tiny structures in therange of 10100 m applying photo-cross-linkable synthetic poly-mers as well as polymerprotein hybrid microstructures. Theydemonstrated freeform fabrication of tubular systems with capil-

lary dimensions using, for example, the above-mentioned PTHF-based materials (Fig.4A andB)[39, 49]. Owing to the hydrophobicnature of the chosen materials, the shape and mechanical proper-ties of the polymer networks remain unchanged in aqueous envir-onments. Once optimized, this may qualify the material toconstitute vascular grafts with improved elasticity, compared withthe PTFE-based or polyester-based grafts that are commonlyused. For application in TE, the generation of porous walls will benecessary in order to enable substance exchange with the

surrounding tissue. By combination of high-resolution TPP withfurther additive techniques such as inkjet printing or stereolithog-raphy, a wide range of structure sizes can be addressed and, thus,this approach is assumed to enable the fabrication of complexgeometries for the mimicking of biological functions.

The Liska laboratory in collaboration with Stamp and collea-gues also investigated photo-elastomers and an additive gener-ation of articial vascular constructs. They presented tubular

structures with 3D microstructured wall architectures with highporosity and high interconnectivity, which they produced bymicrostereolithography out of materials originally presented bySchuster et al.[38,75], thereby demonstrating the potential of themethod (Fig.4C andD). To improve the properties of the materialsin view of articial blood vessels, Baudis et al. developed resinformulations combining diacrylate-mediated cross-linking withdithiol-mediated chain transfer and reactive diluents, based onurethan-diacrylate monomers (Fig. 4E). By use of DLP, they

Figure 4:Complex structures, achieved by application of free-form fabrication methods using light-induced cross-linking of polymers. (Aand B) Scanning electronmicrographs of branched tubular structure, generated by two-photon polymerization of photo-polymerizable ,-polytetrahydrofuranether-diacrylate (PTHF-DA)resins. The height of the tubular structures is 160 m, where the inner diameter and wall thickness are 18 and 3 m, respectively. (Reproduced from Meyeret al.,[49] with the permission of the authors.) (Cand D) SEM images of a highly porous CADCAM structure out of a cytocompatible photopolymer (from [75]), fabricatedby microstereolithography applying a UV laser. (Reproduced from Baudiset al.[38] with the permission of IOP Publishing and the authors.) In both cases, the materialsproved to be cytocompatible, but do not yet meet the required physicochemical properties for articial blood vessels. (E) Graphical illustration of sophisticated mater-ial design. Top: Commonly used hydrogels based on polyethylene glycol diacrylates (PEGDAs) with a high cross-link density along the polymer backbone are sensitiveto cracks as they cannot dissipate energy by inter-chain gliding modes. Middle: The addition of reactive diluents based on methacrylate loosens the network as thecross-link density is reduced. Bottom: The application of chain transfer agents was tested in order to change the polymer architecture and to provide relief to thebrittle material properties that seem to be an intrinsic property of the poly(acrylate) back bone. ( F) Tubular structures fabricated from the new formulations based on(E) using digital light processing (DLP). (Cand D) reproduced from Baudiset al.[38] with the permission of IOP Publishing and the authors). MA: monoacrylate; CTA:chain transfer agent.



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generated polymer networks with reduced cross-linking densityand high contents of reversible H-bonds (Fig. 4F). Based on theresulting gliding of the polymer chains, they produced tough mate-rials with tensile strength (1 MPa), suture tear resistance (300N mm1) and elastic modulus (


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alcohol (PVA). The authors also proved that the generation of cell-laden tubular structures was possible using HeLa cells [52]. Withtheir method, Nakamuraet al. were able to prepare linear tubularstructures with lengths in the centimetre range.

Xu et al. have rened the technique and designed a drop-on-demand inkjet-printing process to vertically fabricate 3D zigzagtube structures [54]. By carefully selecting the operating conditionswith the help of the buoyant force from the CaCl2 solution, they

achieved deposition of 210 layers of Ca-alginate droplets withencapsulated 3T3 broblast cells to form the freestanding tubeshown in Fig.5B with a height of 10 mm, an overhang angle of 63and an overhang height of 5 mm.

These studies show that careful and detailed optimization of thematerials and printing processes is necessary in order to exploitthe potential of printing and microdispensing for advanced TE inthe future. However, alginate is not an ideal material for livingtissue construction as it is not part of the natural ECM of mammalsand does not promote cell adhesion and tissue growth and, thusapoptosis may be induced in encapsulated cells. In addition,though alginate is known to have a high level of bioafnity, it isunknown whether it is biodegradable in a human body. Therefore,the use of more ECM-like materials instead of or along with alginate

gel is required [76].Yet, the stacking of droplets or strands is difcult to control andwithout a stabilizing bath as in the CaCl2-alginate system, ad-equate supporting materials and structures are needed for free-form fabrication processes. To work in combination with thebuilding materials, such support materials also have to full a longcatalogue of requirements, e.g. removability, and the ability toform dened interfaces with building materials, i.e. showingneither repelling effects nor blending. While printing-based tech-niques have the general advantage that different building materi-als can be combined within the process of assemblywhich is notpossible in using cross-linking within a baththe major challengeof the necessity of supporting free-standing structures currentlylimits the resolution in nozzle-based or printer-based techniques.

Visser et al. [60], have tested combinations of various buildingmaterials and support materials in order to improve shape delity ingeneration of complex structures by nozzle-based deposition tech-niques: Hydrogels which are derived from the natural ECM (gelatin,which is derived from collagen and photo-cross-linkable due tomethacrylation (GelMA) together with gellan gum) and Ca-alginateas supportive material; or biodegradable thermoplastic polycapro-lactone (PCL) as building material with PVA as a supportive material.They used piston-driven syringes for hydrogel deposition and screw-driven melt extrusion of PCL and PVA. Removal of PVA supportstructures from PCL was achieved simply by rinsing with water; PCLcould be removed from GelMA gels manually and Ca-alginate couldbe washed away from UV-cured GelMA in sodium citrate solutions.They achieved a free-standing, multi-branched tubular structure fab-

ricated out of PCL strands using PVA support structures in thedimensions of 67 42 8 mm3 (L W H).The open vessel lumenof the branches decreased from 4 mm to 2 mm in diameter. Suchdigitally fabricated constructs are rather stiff and could, for example,serve as degradable scaffolds to be integrated into moulded hydro-gels to temporarily stabilize biomimetic tube structure fortissue-engineering purposes. The GelMA/gellan hydrogel structurespresented by the authors displayed channel diameters of approxi-mately 5 mm and wall thicknesses of 25 mm after removal ofCa-alginate sacricial material. Thus, such bio-based hydrogel ma-terial systems were in the same dimensional range as the perfusablematrix structures described earlier rather than near the aspect ratio

of tubular structures and, thus, demonstrate once again the difcultyof controlling resolution in bottom-up assembly of hydrogels usingdrop-on-demand printing.

Similar materials and methods were applied by Skardal and col-leagues, who prepared cell-laden tubular structures with luminain the upper micrometre to millimetre range (0.55 mm) applyingthe nozzle-based system Fab@Home [5759]. They rst printed acentral, cell-free core disc using a spiral pattern (12 mm), then

added a cell-containing ring around the core (2 mm) and

nallyadded another cell-free supportive outer ring. These steps wereiterated along the z-axis to produce cylindrical structures. In oneapproach, they used chemically modied, photo-cross-linkablegelatin and photo-cross-linkable hyaluronic acid (HA) as buildingmaterials [58]. In another approach, thiol-modied gelatin andthiol-modied HA were cross-linked via multivalent gold nano-particles [57] or tetra-acrylated polyethylene glycol derivatives[59]. In all cases, the building materials were partially cross-linkedprior to printing and further cross-linked after printing, either byUV irradiation or by incubation at 37C. As supportive material,they used non-cross-linkable HA derivatives. HA does not supportcell attachment and thus the non-adhesive surface forced the cellsto remain within the desired tubular structure. To obtain free-

standing tubes with hollow cavities, the supportive HA was thenremoved. Although the aspect ratios achieved were again verylow, this study proved that providing ECM-derived componentswithin the hydrogel matrix supported the viability of the encapsu-lated cells, which were then able to remodel the temporary arti-cial ECM by secretion of endogenous ECM proteins.

Duarte Camposet al.[55] presented an interesting approach incombining high-density uorocarbon as a liquid support materialwith drop-on-demand printing: They generated high-aspect-ratiohollow bres by syringe-based dispensing of agarose hydrogelswith encapsulated human mesenchymal stem cells (hMSCs). Asthe contact angle of agarose gel droplets printed onto the hydro-philic surface of the cross-linked gel was increased within thehydrophobic environment, the precision of the printing was

increased. Histological analyses after 21 days proved cells to beviable with marked matrix production within the constructs. Thus,for the cell types used (hMSCs and human MG-63 cells, an osteo-sarcoma-derived cell line), it is suggested that the polysaccharideagarose served as an adequate temporary matrix and that theuorocarbon support did not signicantly harm the encapsulatedcells.

Another aspect in view of future reconstruction of native bloodvessels was captured by Kucukgul et al. [77], who have capturedthe geometry of an aorta from magnetic resonance imaging dataand have then segmented and converted them into 3D computermodels. Yet, due to the above-mentioned difculties in stabilizingtheir constructs, the authors have presented the proof of theirdata conversion by an eight-layer model only.

Very straightforward approaches also introduced direct extrusionof hollow bres using coaxial nozzles. While this method cannotprovide any branched tubular structures, perfusion systems for 3Dcell culture or even fabrication of small-diameter vascular grafts formedical applications may be achieved in the future. Luoet al. fromthe Gelinsky laboratory have applied the Ca-alginate system for theproduction of stacks of inter-bred hollowbres [78]. They generatedfreestanding hollowbres by continuous extrusion of highly concen-trated alginate (16.7% w/w) with PVA (6% w/v) solution using ahome-built shell/core nozzle tip. After extrusion, the scaffolds weretransferred to CaCl2 solution in order to achieve the complexionof the alginate and dissolution of the PVA. The outer and inner



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diameters of the hollowbres were adjusted to 3001000 and 150450 m, respectively, and multilayer stacks were prepared (Fig. 6AandB).

A similar approach was introduced by Zhang, Yu and Ozbolat(Fig.6CandD). They fabricated hollow tubes without support bydispensing less concentrated (26%), cell-laden sodium alginatethrough the outer tube of a coaxial nozzle system with simultan-eously ushing of CaCl2 cross-linker solution through the innertube [6264]. They also tested chitosan and sodium hydroxide so-lution with less success. In their study, the authors used cartilageprogenitor cells and reported that, depending on process and ma-terial parameters such as nozzle diameters, dispensing pressureand alginate concentration, the viability of the encapsulated cellsdropped signicantly at early time points after printing but recov-ered over the time course of 72 h. This suggests that, for the specif-ic cell typecartilage cells generally display spherical morphology,without distinct adhesion to the surrounding matrixalginate-

based matrices might act as adequate matrix material.As all scaffold-based biomaterials have to deal with challenges

such as cyto- and biocompatibility, degradation behaviour orpotential immunogenicity, Norotte and Jakab from the Forgacs la-boratory, which is also the origin of the bioprinting, and bioengin-eering companies Organovo (Organovo Holdings, Inc., CA, USA)and Modern Meadow (Modern Meadow LLC, Columbia, MO,USA), nally focus on fabrication of scaffold-free tissue constructsbased only on cells and the genuine matrix they secrete. Theyapplied one-by-one bioplotting of cell aggregates consisting ofeither one cell type or several cell types for the fabrication oftubular structures [61,71]. In a rst approach, vertical tubes were

fabricated by alternate deposition of the spherical cell aggregatesand collagen, which served as sacricial supporting material. Thecell spheroids fused within 57 days into stable tubular structures

and, thus, after removal of the stabilizing collagen sheets, hollowtubes were achieved. One limitation of this early approach wasthat the supportive collagen layers could not be completelyremoved since cells integrated the collagen into thenal structure.Thus, in a second approach, they replaced collagen with agarose,which is biologically inert and cannot be invaded or rearrangedby cells. Moreover, they found that, instead of tissue spheroids,the use of multicellular cylinders represented considerable pro-gress in terms of time and precision of the nal structure. By com-bination of multicellular cylinders made from HUVSMCs andhuman dermal broblasts (HDFs), double-layered vascular wallscould be built. The fused articial vessels were sufciently sturdyto be handled and transferred into specically designed bioreac-tors for further maturation. In these studies, the smallest tube

achieved was 90 m in diameter and obtained a wall thickness of300 m. This approach is particularly interesting because vesselwalls based on the modelling activities of the cells are achieved.

CONCLUSION

The generation of vascularized tissue is one of the key challengesin TE. The evaluation of free-form fabrication methods for thegeneration of perfusable tissue is an important step towards thisdemanding goal.

Figure 6:Plotting of hollowbres by the use of core/shell nozzles directly provided vessel-like perfusable tubes for developing embedded channels serving as a vascu-lar network for thick tissue fabrication. (Aand B) plotted alginate/polyvinyl alcohol scaffold consisting of layers of hollow strands. (Reproduced from Luo et al. [78],with permission of John Wiley and Sons.) (C) Media perfusion through an alginate microuidic channel, 44 cm in length and 1 mm in width with 500-m lumen diam-eter. (D) Cartilage cells were encapsulated in the hydrogel solution during the extrusion process (green uorescence: viable cells, red: dead cells). (Reproduced fromZhanget al.[64] with permission of the American Society of Mechanical Engineers (ASME).).



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From our current point of view, the fabrication of perfusablechannel systems within cell-laden hydrogel matrices has provedto be the most straightforward method to generate large 3D tissueconstructs by enabling the exchange of O2and CO2, and nutrientsupply within the 3D matrix bulk. This approach allows fulllmentof the initially recognized requirements of both, dened inlet andoutlet branches connected to a channel network with biomimeticgeometry for undisturbed uid dynamics, and the conditions for

formation of new capillaries through angiogenesis. Thus, we expectthis approach to be the crucial next step that will initiate real 3D TE.First proofs of concept showed that cartridge-based drop-on-

demand techniques can be expected to be of great use for thecombination of different matrices and cells in a 3D organizedmanner as required for biomimetic assembly of tissue. Yet, theyseem to be limited for assembly of freestanding high-aspect-ratiostructures.

Extrusion of hollow bres does not provide the full range ofdegrees of freedom that are connected to free-form fabricationmethods, yet we appreciate the method to be straightforward forpreparation of linear channels and also for basic supply of 3Dtissue models. Additionally, the presented approaches stronglyfeed our vision of completely new generation of vessel substitu-

tions for cardiovascular diseases. Yet, numerous challenges haveto be met and very accurate developments in materials scienceare needed, whether based on synthetic or biological polymers, inorder to match the mechanical and biological properties of nativevessels such as compliance, tear strength, bursting strength, hae-mocompatibility and long-term stability. The full structure andfunction may even be established by integrated cells remodellingthe bioarticial vessel wall. Thus, we assume that 510 years of re-search and development are additionally needed to achieve thedemanding goal of bioprinting blood vessels for the rst in vivotrials.

ACKNOWLEDGEMENTS

The authors thank the Fraunhofer Gesellschaft, the EuropeanCommission (Artivasc 3D GA 263 416), and the Peter und TraudlEngelhorn Stiftung for nancial support.

Conict of interest: none declared.

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